Quantification of optical absorption coefficients using acoustic spectra in photoacoustic tomography

ABSTRACT

Accurately quantifying optical absorption coefficient using acoustic spectra of photoacoustic signals. Optical absorption is closely associated with many physiological parameters, such as the concentration and oxygen saturation of hemoglobin, and it can be used to quantify the concentrations of non-fluorescent molecules. A sample is illuminated by, for example, a pulsed laser and following the absorption of optical energy, a photoacoustic pressure is generated via thermo-elastic expansion. The acoustic waves then propagate and are detected by a transducer. The optical absorption coefficient of the sample is quantified from spectra of the measured photoacoustic signals. Factors, such as system bandwidth and acoustic attenuation, may affect the quantification but are canceled by dividing the acoustic spectra measured at multiple optical wavelengths.

CROSS REFERENCE TO RELATED APPLICATIONS

This application is a continuation of U.S. application Ser. No.13/637,897, filed Oct. 24, 2012 which was the National Stage ofInternational Application No. PCT/US2011/031823, filed Apr. 8, 2011,which claims the benefit of U.S. Provisional Patent Application No.61/322,605 filed Apr. 9, 2010, all of which are hereby incorporated byreference in their entirety.

STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH & DEVELOPMENT

This invention was made with government support under grant R01EB008085, awarded by the U.S. National Institutes of Health. Thegovernment has certain rights in the invention.

BACKGROUND

The embodiments described herein relate generally to photoacousticscanning methods and apparatus and, more particularly, to a method and asystem for determining an optical absorption coefficient by an objectusing acoustic spectra.

Total, oxygenated, and deoxygenated hemoglobin concentrations ([HbT],[HbO₂], and [HbR], respectively) are fundamental pathophysiologicalparameters in biomedicine. For example, abnormally low [HbT] may becaused by loss of blood, nutritional deficiency, chemotherapy,inflammation, kidney failure or bone marrow problems, while abnormallyhigh [HbT] may be related to exposure to high altitude, smoking,dehydration and tumors. Blood oxygen saturation (sO₂), which is definedas [HbO₂] divided by [HbT], is vital in understanding brain hemodynamicsin response to sensory stimulations, monitoring healing of burns andwounds, and evaluating the effectiveness of chemotherapy andradiotherapy on tumors. Several techniques have been developed toquantify hemoglobin concentration and sO₂ in vivo, includingnear-infrared spectroscopy (NIRS), blood oxygen level dependent (BOLD)contrast magnetic resonance imaging (MRI), electron paramagneticresonance imaging (EPRI), positron emission tomography (PET), and singlephoton emission computed tomography (SPECT). However, all of thesemodalities have disadvantages. For example, at least some of thesemodalities have poor spatial resolution, relative quantification, andundesirable contrast agent injection. Photoacoustic (PA) tomography(PAT) has already demonstrated its ability to monitor biologicalhemodynamic functions without using exogenous contrast agents.Quantitative PAT is challenging, because compensating for the fluence inquantitative in vivo photoacoustic tomography is difficult, and factors,such as the tissue acoustic attenuation and the imaging systembandwidth, also affect the quantification accuracy.

BRIEF DESCRIPTION

In one aspect, a photoacoustic imaging method includes illuminating anobject using a light beam emitted by a light source, and detecting apressure wave emitted by the object using an acoustic transducer,wherein the pressure wave is induced by the object in response to thelight beam (e.g., light pulse). The object is illuminated with at leasttwo light pulses having different wavelengths, and the absorptioncoefficient of the object is determined from the detected pressurewaves. The method also includes calculating an amount of optical energythat is absorbed by the object from the light beam, based at least onthe pressure.

BRIEF DESCRIPTION OF THE DRAWINGS

The embodiments described herein may be better understood by referringto the following description in conjunction with the accompanyingdrawings.

FIG. 1 is a schematic diagram of an exemplary optical resolutionphotoacoustic microscopy (OR-PAM) system.

FIG. 2 is a diagram of an exemplary experimental setup of a sample forOR-PAM.

FIG. 3 is a set of graphs that illustrate an example of an objectspectrum O(ωλ), a system dependent response H(ω), and a tissue-relatedacoustic attenuation effect a(ω) acquired via the OR-PAM system of FIGS.1 and 2.

FIG. 4 is a graph illustrating acoustic spectra of a number ofphotoacoustic signals.

FIG. 5 is a graph illustrating a comparison of photoacoustic spectra oftwo objects.

FIG. 6 is a graph illustrating quantified absorption coefficients of aplurality of samples.

FIG. 7 is a graph illustrating acoustic spectra of a photoacousticsignal with and without an optical phantom layer.

FIG. 8 is a photoacoustic maximum amplitude projection image of a bloodvessel.

FIG. 9 is a graph that illustrates a ratio of acoustic spectra at twodifferent optical wavelengths.

FIG. 10 is a diagram that illustrates raster scanning an acoustictransducer to obtain three-dimensional images via photoacousticmicroscopy.

FIG. 11 is a diagram that illustrates light diffusion at or near a topsurface of a blood vessel.

FIG. 12 is a diagram that illustrates a characterization of light on thetop surface of the blood vessel of FIG. 11 as an isotropic point source.

FIG. 13 is a diagram that illustrates an alternative embodiment of anOR-PAM that includes a plurality of acoustic transducers arranged intoan array.

FIG. 14 is a flow chart of a method of determining an absorptioncoefficient of a volume of an object.

DETAILED DESCRIPTION

While the making and using of various embodiments of the presentdisclosure are discussed in detail below, it should be appreciated thatthe present disclosure provides many applicable inventive concepts thatcan be embodied in a wide variety of specific contexts. The specificembodiments discussed herein are merely illustrative of specific ways tomake and use the disclosure and do not delimit the scope of thedisclosure.

To facilitate the understanding of the embodiments described herein, anumber of terms are defined below. The terms defined herein havemeanings as commonly understood by a person of ordinary skill in theareas relevant to the present disclosure. Terms such as “a,” “an,” and“the” are not intended to refer to only a singular entity, but ratherinclude the general class of which a specific example may be used forillustration. The terminology herein is used to describe specificembodiments of the disclosure, but their usage does not delimit thedisclosure, except as outlined in the claims.

To be consistent with the commonly used terminology, whenever possible,the terms used herein will follow the definitions recommended by theOptical Society of America (OCIS codes).

In some embodiments, the term “photoacoustic microscopy” refersgenerally to a photoacoustic imaging technology that detects pressurewaves generated by light absorption in the volume of a material (such asbiological tissue) and propagated to the surface of the material. Inother words, photoacoustic microscopy is a method for obtainingthree-dimensional images of the optical contrast of a material bydetecting acoustic or pressure waves traveling from the object. Imagingresolution is on the micrometer scale.

In some embodiments, the term “photoacoustic tomography” also refers toa photoacoustic imaging technology that detects acoustic or pressurewaves generated by light absorption in the volume of a material (such asbiological tissue) and propagated to the surface of the material. Theemphasis is sometimes on photoacoustic computed tomography, i.e.,cross-sectional or three-dimensional photoacoustic imaging based oncomputer reconstruction, although the most general definition ofphotoacoustic tomography encompasses photoacoustic microscopy.

In some embodiments, the terms “reflection mode” and “transmission mode”refer generally to a laser photoacoustic microscopy system that employsthe detection of acoustic or pressure waves transmitted from the volumeof their generation to the optically irradiated surface and a surfacethat is opposite to, or substantially different from, the irradiatedsurface, respectively.

In some embodiments, the term “time-resolved detection” refers generallyto the recording of the time history of a pressure wave with a temporalresolution sufficient to reconstruct the pressure wave profile.

In some embodiments, the term “photoacoustic waves” refers generally topressure waves produced by light absorption.

In photoacoustic imaging, a sample is illuminated, usually by a pulsedlaser, and following the absorption of optical energy, an initialpressure is generated via thermo-elastic expansion. The photoacousticwaves then propagate and are detected by acoustic or ultrasonic sensors.The strength of the initial pressure is directly proportional to theabsorbed optical energy in the tissue and, therefore, multi-wavelengthphotoacoustic measurements may yield optical absorption spectralinformation. Since ultrasound scattering is much weaker than opticalscattering in biological tissues, photoacoustic tomography is capable ofhigh resolution imaging at depths beyond the optical transport mean freepath. Moreover, because photoacoustic imaging inherently exploitsoptical absorption contrast, it usually has a higher sensitivity tooptical absorption than other optical imaging technologies. By measuringthe optical absorption spectrum with photoacoustic imaging, sO₂ may bequantified in the same way near infrared spectroscopy (NIRS) quantifiessO₂, except with high spatial resolution and sensitivity.

Photoacoustic images are spatial mappings of the absorbed optical energyA( r) (Jm⁻³), which is the product of the optical absorption coefficientμ_(a)( r, λ) (m ⁻¹) and the fluence F( r) (Jm⁻²). To obtain theintrinsic tissue property μ_(a)( r, λ), there is a need to compensatefor the extrinsic quantity F( r). Since the optical parameters fortissues are usually heterogeneous and unknown, the fluence varies fromcase to case and is difficult to model. As a consequence, compensatingfor the fluence in quantitative in vivo photoacoustic tomography is achallenge.

Currently, fluence compensation can be done invasively ornon-invasively. The invasive method includes positioning an opticalabsorber with a known spectrum close to the region of interest, and thennormalizing the measured photoacoustic signals from the objects with theamplitudes of the photoacoustic signals from the optical absorber havinga known spectrum. Alternatively, the incident fluence may be estimatedby measuring light transmission through a sample of excised tissue ofthe same type as the tissue overlaying a region of interest. Thenon-invasive method involves solving the radiative transfer equation(RTE) and photoacoustic wave equations with iterative algorithms. Bothof these methods are based on the linear relationship between the localfluence incident on the blood vessel, acquired either by experimentalmeasurements or numerical simulation, and the peak amplitude of thephotoacoustic signal.

A temporal profile of the photoacoustic signal is used to quantify theoptical absorption coefficients. This temporal method isself-calibrating since it depends on the relative temporal profilerather than the absolute amplitude of the photoacoustic signal.Therefore, the temporal profile is less dependent on changes in theoptical properties of overlying tissues. However, since the temporalprofile is distorted by various factors, such as the limited bandwidthof acoustic detectors and frequency-dependent acoustic attenuation ofthe sample or region of interest, linear translation of the temporalprofile to the optical absorption may be inaccurate.

As will be described in greater detail below, embodiments of the presentdisclosure provide a method for quantifying the optical absorptioncoefficient using the acoustic spectra of the photoacoustic signals. Bycarefully investigating the factors involved in generating the acousticspectrum, the effects of detector bandwidth and acoustic attenuation areeliminated, as shown later. At least some embodiments of the method areself-calibrating since it deals only with the relative change in variousacoustic frequencies. The acoustic spectrum S(ω) of the receivedphotoacoustic signal depends on three factors: 1) the ‘real’ objectspectrum O(ω, λ) measured with unit fluence, which is related to thetarget object's shape, size, optical properties, and fluence incidentdirections; 2) the system dependent response H(ω), which is the Fouriertransform of the photoacoustic signal from an ideal point absorbermeasured with this system without acoustic attenuation in the tissue;and 3) the tissue related acoustic attenuation effect a(ω), which isrelated to the acoustic properties of the tissue that lies between thetarget objects (e.g., region of interest) and the detector. Based on asystem linearity assumption, the above factors lead to the expressionS(ω, λ)=F(λ)O(ω, λ)H(ω)a(ω). The last two terms remain unchanged whensamples are measured with the same system under the same condition, andtherefore are usually cancellable. An example is where light at variousoptical wavelengths is used to excite one blood vessel. Therefore, bydividing the photoacoustic acoustic spectrum measured at one opticalwavelength by the spectrum measured at another wavelength, the systemdependent effects and the acoustic attenuation effect may be eliminated.As such, the absolute value of μ_(a) may be quantified with this methodeven though F varies with the optical wavelength. By contrast, previousmethods can quantify only the relative value of μ_(a).

As a feasibility study, this idea has been implemented and validatedusing one form of photoacoustic imaging, optical resolutionphotoacoustic microscopy (OR-PAM), where the object spectrum O(ω, λ) maybe relatively easily modeled. Referring to FIG. 1, in OR-PAM,photoacoustic A-scan signals are acquired through time-resolved acousticdetection, and three-dimensional images are formed by raster scanningthe acoustic transducer 102 along the transverse plane, as shown inFIG. 1. The axial resolution of the system depended on the transducerbandwidth (centered at 50 MHz with 80% bandwidth), while the lateralresolution relied on optical focusing, which can reach the theoreticaloptical diffraction limit. For the system of FIG. 1, the axial andlateral resolutions were quantified to be approximately 15 micrometers(μm) and approximately 5 μm, respectively. Therefore, the surface ofblood vessels with a diameter of greater than approximately 30 μm may beroughly treated as a flat surface. In this case, the acoustic spectrumof the generated photoacoustic signal is only related to the opticalpenetration depth. If F₀ is used to denote the incident fluence on thesurface of the blood vessel, the fluence inside the blood vessel obeysBeer's law and may be written as F(z)=F₀ exp(−μ_(a)z), as shown in FIG.2. Here, the reduced scattering coefficient is much less than theabsorption coefficient, because the anisotropy factor is so close to 1in blood in the optical spectral region that was used (around 585 nm);therefore, it is neglected. The photoacoustic signal generated by theobject is expressed using Equation (1):

O(t, λ)=μ_(a)(λ)exp [−μ _(a)(λ)ct]  Eq. (1)

where c is the speed of sound in the biological tissue. Performing aFourier transformation of Eq. (1) leads to Equation (2):

$\begin{matrix}{{{O\left( {\omega,\lambda} \right)}} = \frac{1}{\sqrt{\left\lbrack {\omega/{\mu_{a}(\lambda)}} \right\rbrack^{2} + c^{2}}}} & {{Eq}.\mspace{14mu} (2)}\end{matrix}$

Moreover, if the photoacoustic signals of the blood vessel are measuredat two optical wavelengths, the ratio of the spectra of thephotoacoustic signals may be written as shown in Equation (3):

$\begin{matrix}\begin{matrix}{\frac{S\left( {\omega,\lambda_{1}} \right)}{S\left( {\omega,\lambda_{2}} \right)} = \frac{{F\left( \lambda_{1} \right)}{O\left( {\omega,\lambda_{1}} \right)}{H(\omega)}{a(\omega)}}{{F\left( \lambda_{2} \right)}{O\left( {\omega,\lambda_{2}} \right)}{H(\omega)}{a(\omega)}}} \\{= \frac{{F\left( \lambda_{1} \right)}\sqrt{\left\lbrack {\omega/{\mu_{a}\left( \lambda_{2\;} \right)}} \right\rbrack^{2} + c^{2}}}{{F\left( \lambda_{2} \right)}\sqrt{\left\lbrack {\omega/{\mu_{a}\left( \lambda_{1} \right)}} \right\rbrack^{2} + c^{2}}}}\end{matrix} & {{Eq}.\mspace{14mu} (3)}\end{matrix}$

Furthermore, the absolute values of μ_(a)(λ₁), μ_(a)(λ₂) and F(λ₁)/F(λ₂)may be derived by fitting the ratio of Eq. (3). The assumption of laserbeam collimation within the blood vessel for Eqs. (1)-(3) is validwithin and far from the optical focal zone. Otherwise, the 3D Gaussianbeam profile should be considered. Here, the Rayleigh range of theGaussian beam is approximately 30 μm, which is comparable to the opticalpenetration depth in blood around 585 nm.

FIG. 1 is a schematic diagram of an exemplary OR-PAM system, FIG. 2 is adiagram of an exemplary experimental setup, and FIG. 3 is a set ofgraphs that illustrate an example of O(ω, λ), H(ω), and a(ω). The systemincludes a dye laser 104 pumped by a Nd: YLF laser 106 that is used asthe irradiation source. The laser beam from the dye laser 104 passesthrough a condenser lens 140, is spatially filtered by a pinhole 108, isredirected by a mirror 142, and is then focused by an objective lens110. Sonic and ultrasonic focusing is achieved through a plano-concavelens 112. The optical objective lens 110 and 50 MHz acoustic transducer102 are confocally configured via a correction lens 122, a right angleprism 124, and a silicone oil layer 126. A water tank 128 and membrane130 separate the acoustic lens 112 from the sample 132. Volumetricimages are generated by data acquisition device 114 through acombination of time-resolved detection of the photoacoustic waves with atwo-dimensional raster scanning in the transverse plane. A scannercontroller 116 and scanner 118 (i.e., two dimensional motivatorplatform) provide rasterizing under the control of data acquisitiondevice 114, and the data acquisition device 114 receives data fromacoustic transducer 102 via amplifier 120. The data acquisition device114 provides a trigger signal to the pump laser 106 and the scannercontroller 116 and also provides a clock signal to the scannercontroller 116. The trigger and clock signals initiate and control thetiming of light pulses from the pump laser 106 and dye laser 104 incoordination with movement of the scanner 118.

As shown in FIG. 2, the optical focus 202 of the system of FIG. 1 ismuch smaller than a targeted blood vessel whose top surface 204 withinthe optical focal diameter can therefore be approximated as a plane. Anoverlying tissue 206 covers the blood vessel top surface 204, and blood208 lies beneath the blood vessel top surface 204 in the setup of FIGS.1 and 2. The optical fluence within the blood vessel decaysexponentially with depth at a rate of the optical absorptioncoefficient. FIG. 3 is an example of the object spectrum O(ω, λ), thesystem dependent response H(ω), and the tissue related acousticattenuation effect a(ω).

In a phantom study, original black ink was diluted with water in sixratios ranging from 1:1 to 1:6. The original and diluted ink sampleswere sequentially placed in a container, sealed with plastic membrane,and then the container was placed in a water tank. Photoacoustic A-linesignals were acquired from these samples. The acoustic spectra of thephotoacoustic signals are shown in FIG. 4. Compared with the spectrum ofthe photoacoustic signal from the original ink sample, the spectra ofthe photoacoustic signals from the diluted ink samples are “shifted” tolower frequencies. Light penetrated deeper in lower concentration inksamples, and the corresponding photoacoustic signals decay more slowlyin the time domain. Therefore, the spectra contain more low-frequencycomponents. By dividing the measured spectra of any two ink samplesfrequency-by-frequency, as shown in FIG. 5, the absorption coefficientsof both samples may be determined by fitting the resultant ratio curvewith Eq. (3). Because parts of the spectra (e.g., 0-5 MHz and greaterthan 70 MHz in the exemplary system) are unreliable due to thelimited-band detection, they are not used for the fitting. Thequantified absorption coefficients of all seven samples and theirtheoretical values are plotted in FIG. 6. To demonstrate whether therecovered absorption coefficients are independent of acousticattenuation and optical fluence, three of the ink samples were coveredwith an identical layer of optical phantom (˜1.5 mm 2% Agar, 0.1%intralipid, 1% black ink). The spectra of the photoacoustic signals fromone ink sample with and without this layer are shown in FIG. 7. Thespectral profiles differ because of the acoustic attenuation, while thespectral magnitudes differ owing to the optical fluence attenuation.Since the acoustic properties of the layer added between the samples andthe detector are the same for the three ink samples, the acousticattenuation may be cancelled by taking the ratio of the acoustic spectraof photoacoustic signals from any two covered ink samples. Thequantified absorption coefficients of these samples agree with theexpected values as shown in FIG. 6.

In an in vivo experiment, a region measuring approximately 1 squaremillimeter (mm²) in a nude mouse ear was imaged with two opticalwavelengths of 561 nanometers (nm) and 570 nm. FIG. 8 shows thephotoacoustic maximum amplitude projection (MAP) image acquired with anoptical wavelength of 570 nm, which is an oxygen insensitive absorptionwavelength of hemoglobin. Each point in the MAP image records themaximum value of a Hilbert transformed photoacoustic A-scan. Two vesselsmarked with V₁ and V₂ in FIG. 8 were selected for a quantitative study.The A-scans acquired within these two vessels were properly aligned andthen averaged. For each vessel, the acoustic spectrum measured at 561 nmwas divided point by point by the acoustic spectrum measured at 570 nm,and the absorption coefficients were acquired by fitting the ratio withEq. (3), as shown in FIG. 9. The [HbT], [HbO₂], and [HbR], together withthe sO₂ values were calculated based on the quantified opticalabsorption coefficients at the two optical wavelengths, as shown belowin Eqs.(4) and (5) and Table 1.

$\begin{matrix}{\begin{bmatrix}\left\lbrack {HbO}_{2} \right\rbrack \\\lbrack{HbR}\rbrack\end{bmatrix} = {{\frac{64500}{2.303}\begin{bmatrix}{ɛ_{ox}\left( \lambda_{1} \right)} & {ɛ_{de}\left( \lambda_{1} \right)} \\{ɛ_{ox}\left( \lambda_{2} \right)} & {ɛ_{de}\left( \lambda_{2} \right)}\end{bmatrix}}^{- 1}\begin{bmatrix}{\mu_{a}\left( \lambda_{1} \right)} \\{\mu_{a\;}\left( \lambda_{2} \right)}\end{bmatrix}}} & {{Eq}.\mspace{14mu} (4)} \\{{sO}_{2} = \frac{\left\lbrack {HbO}_{2} \right\rbrack}{\left\lbrack {HbO}_{2} \right\rbrack + \lbrack{HbR}\rbrack}} & {{Eq}.\mspace{14mu} (5)}\end{matrix}$

In Eq.(4), λ₁ and λ₂ are the two wavelengths, and ε_(ox) and ε_(de) arethe known molar extinction coefficients of oxy- and deoxyhemoglobin,respectively. According to the sO₂ values, V₁ and V₂ were identified tobe an arteriole-venule pair. The incident fluence ratio at the twooptical wavelengths F(λ₁)/F(λ₂) was also quantified for both vessels. Inthis special case, the two vessels are embedded at a similar depth, andthe optical and the acoustic properties of the overlying tissue arecomparable. Here, the quantified fluence ratios turned out to be thesame for V₁ and V₂. If the wavelength-dependent fluence variations areignored by simply assuming F(λ₁)/F(λ₂)=1, as shown in the last column ofTable 1, the quantified sO₂ values become inaccurate by approximately 8%and 11% for the artery and the vein, respectively.

TABLE 1 μ_(a)(λ₁) μ_(a)(λ₂) [HbT] [HbO₂] [HbR] sO₂ (cm⁻¹) (cm⁻¹)F(λ₂)/F(λ₁) (g/L) (g/L) (g/L) sO₂ F₂/F₁ = 1 V₁ 143 ± 3 188 ± 4 0.96 ±0.01 110.6 ± 8.1 106.2 ± 4.3  4.4 ± 3.8 0.96 ± 0.04 0.88 (Artery) V₂ 159± 4 186 ± 5 0.96 ± 0.01 110.2 ± 9.2  77.1 ± 4.9 33.1 ± 4.3 0.70 ± 0.070.62 (Vein)

Embodiments of the method may be applied in the optical diffusive regimewith photoacoustic microscopy, whose lateral resolution relies on theacoustic focus (e.g., ˜50 μm at 50 MHz acoustic frequency). As shown inFIG. 10, photoacoustic A-scan signals are acquired through time-resolvedacoustic detection, and three-dimensional images are formed by rasterscanning an acoustic transducer 402 and optical components (e.g.,optical illumination source 404 and mirror 406) along a transverse plane(e.g., transverse to the direction of the optical illumination strikinga surface 408 of a sample 410). In the embodiment of FIG. 10, thesurface 408 of the sample 410 is illuminated with an annular ring havinga dark center 412 as shown in the call out from the main figure. Thesurface of blood vessels with sufficiently large diameter (e.g., greaterthan 300 μm for the 50-μm lateral resolution) may be approximatelytreated as a flat surface. As shown in FIG. 11, it may generally beassumed that the light is completely diffused (e.g., see diffused light504) when it reaches a top surface 502 of a blood vessel within tissue410. Blood 506 is located below the blood vessel top surface 502,overlying tissue 508 is located above the blood vessel top surface 502,and the acoustic focus 510 narrows at the surface of the blood vessel502 as shown in FIG. 11. Moreover, as shown in FIG. 12, light at eachpoint (e.g., point 602) on the top surface of the blood vessel 502 canbe seen as an isotropic point source for each point (e.g., point 604) ina thin layer of blood 606, and the fluence F(z) in the blood vessel canbe expressed as shown in Equation (6):

$\begin{matrix}\begin{matrix}{{F(z)} = {\pi \; R^{2}{\int{\int_{S^{\prime}}{{\exp \left( {{- \mu_{a}}\sqrt{x^{\prime 2} + y^{\prime 2} + z^{\prime 2}}} \right)}{x^{\prime}}{y^{\prime}}}}}}} \\{= {F_{0}\pi^{2}R^{2}{\int_{0}^{\infty}{{\exp \left( {{- \mu_{a}}\sqrt{u + z^{2}}} \right)}\ {u}}}}}\end{matrix} & {{Eq}.\mspace{14mu} (6)}\end{matrix}$

where R is the radius of the acoustic focus 510. The object spectrumO(ω, λ) may then be calculated using Fourier transformation, where z isfirst converted to time t through z=ct. Similar to the case of OR-PAM,the system-dependent response H(ω) and the tissue related acousticattenuation effect a(ω) are canceled by dividing, at each acousticfrequency, the acoustic spectra measured at two optical wavelengths.FIG. 13 shows another embodiment of this method with an acoustictransducer array 702. With a synthetic aperture focusing technique(SAFT), a virtual acoustic focus704 may be formed at any position in thefield of view by applying proper time delays to each element of thearray system. FIG. 13 includes diffused light 706 on the surface ofblood vessel 710 within tissue 708. Using the same process as describedabove, the optical absorption coefficients may be extracted from theacoustic spectra measured at multiple optical wavelengths. Although alinear array is illustrated in the figure, arrays of other shapes—suchas a semicircle, a hemisphere, and a 2D plane—can be used as well.

To increase the accuracy of this method, an acoustic transducer with anappropriate bandwidth is selected. The acoustic spectrum of thephotoacoustic signal is related to the light penetration depth.Therefore, the central frequency of the transducer should vary with thepenetration depth to achieve the best signal to noise ratio (SNR).Therefore, the central frequency of the transducer should match thepenetration depth to maximize signal to noise ratio (SNR). Moreover, SNRis usually low at high acoustic frequencies due to acoustic attenuation.O(ω, λ), H(ω), and a(ω) are all band-limited, and in an exemplaryembodiment, H(ω) is chosen to match O(ω, λ) and a(ω).

Referring to FIG. 14, a method of determining an absorption coefficientof a volume of an object begins with a focusing element (e.g., anobjective or lens) focuses a first light pulse from a light source(e.g., a laser) into a volume of an object or sample at 802. At 804, atransducer detects a first acoustic spectrum emitted by the object inresponse to the volume of the object receiving the first light pulse. At806, the focusing element focuses a second light pulse from the lightsource into the volume of the object, and at 808, the transducer detectsa second acoustic spectrum emitted by the object in response to thevolume of the object receiving the second light pulse. In operation, thetransducer receives an acoustic wave from the object and performs aFourier transform to generate the acoustic spectrum. It is contemplatedthat the Fourier transform may be performed by a component separate fromthe transducer component receiving the acoustic waveform. At 810, acontroller determines an absorption coefficient from the detected firstand second acoustic spectra by dividing the first acoustic spectrum bythe second acoustic spectrum frequency by frequency yielding anormalized acoustic spectrum. The normalized acoustic spectrum is fittedto an ideal curve of an absorption coefficient to determine theabsorption coefficient of the volume of the object. The first and secondlight pulses have different wavelengths, and the method may be repeatedfor multiple volumes of the object and to obtain an image of the object.

It will be understood that the particular embodiments described hereinare shown by way of illustration and not as limitations of thedisclosure. The principal features of this disclosure may be employed invarious embodiments without departing from the scope of the disclosure.Those of ordinary skill in the art will recognize numerous equivalentsto the specific procedures described herein. Such equivalents areconsidered to be within the scope of this disclosure and are covered bythe claims.

All of the compositions and/or methods disclosed and claimed herein maybe made and/or executed without undue experimentation in light of thepresent disclosure. While the compositions and methods of thisdisclosure have been described in terms of the embodiments includedherein, it will be apparent to those of ordinary skill in the art thatvariations may be applied to the compositions and/or methods and in thesteps or in the sequence of steps of the method described herein withoutdeparting from the concept, spirit, and scope of the disclosure. Allsuch similar substitutes and modifications apparent to those skilled inthe art are deemed to be within the spirit, scope, and concept of thedisclosure as defined by the appended claims.

It will be understood by those of skill in the art that information andsignals may be represented using any of a variety of differenttechnologies and techniques (e.g., data, instructions, commands,information, signals, bits, symbols, and chips may be represented byvoltages, currents, electromagnetic waves, magnetic fields or particles,optical fields or particles, or any combination thereof). Likewise, thevarious illustrative logical blocks, modules, circuits, and algorithmsteps described herein may be implemented as electronic hardware,computer software, or combinations of both, depending on the applicationand functionality. Moreover, the various logical blocks, modules, andcircuits described herein may be implemented or performed with a generalpurpose processor (e.g., microprocessor, conventional processor,controller, microcontroller, state machine or combination of computingdevices), a digital signal processor (“DSP”), an application specificintegrated circuit (“ASIC”), a field programmable gate array (“FPGA”) orother programmable logic device, discrete gate or transistor logic,discrete hardware components, or any combination thereof designed toperform the functions described herein. Similarly, steps of a method orprocess described herein may be embodied directly in hardware, in asoftware module executed by a processor, or in a combination of the two.A software module may reside in RAM memory, flash memory, ROM memory,EPROM memory, EEPROM memory, registers, hard disk, a removable disk, aCD-ROM, or any other form of storage medium known in the art. Althoughpreferred embodiments of the present disclosure have been described indetail, it will be understood by those skilled in the art that variousmodifications can be made therein without departing from the spirit andscope of the disclosure as set forth in the appended claims.

A controller, computing device, or computer, such as described herein,includes at least one or more processors or processing units and asystem memory. The controller typically also includes at least some formof computer readable media. By way of example and not limitation,computer readable media may include computer storage media andcommunication media. Computer storage media may include volatile andnonvolatile, removable and non-removable media implemented in any methodor technology that enables storage of information, such as computerreadable instructions, data structures, program modules, or other data.Communication media typically embody computer readable instructions,data structures, program modules, or other data in a modulated datasignal such as a carrier wave or other transport mechanism and includeany information delivery media. Those skilled in the art should befamiliar with the modulated data signal, which has one or more of itscharacteristics set or changed in such a manner as to encode informationin the signal. Combinations of any of the above are also included withinthe scope of computer readable media.

This written description uses examples to disclose the disclosure,including the best mode, and also to enable any person skilled in theart to practice the disclosure, including making and using any devicesor systems and performing any incorporated methods. The patentable scopeof the disclosure is defined by the claims, and may include otherexamples that occur to those skilled in the art. Such other examples areintended to be within the scope of the claims if they have structuralelements that do not differ from the literal language of the claims, orif they include equivalent structural elements with insubstantialdifferences from the literal language of the claims.

What is claimed is:
 1. A method of quantifying an optical absorptioncoefficient of an object, said method comprising: directing a firstlight pulse emitted by a light source into a volume of the object,wherein the first light pulse has a first wavelength; receiving at anacoustic transducer a first acoustic wave emitted by the object inresponse to receiving the first light pulse; performing on a controllera Fourier transformation on the received first acoustic wave to generatea first acoustic spectrum; directing a second light pulse emitted by thelight source into the volume of the object, wherein the second lightpulse has a second wavelength; receiving at the acoustic transducer asecond acoustic wave emitted by the object in response to receiving thesecond light pulse; performing on the controller a Fouriertransformation on the received second acoustic wave to generate a secondacoustic spectrum; and quantifying the absorption coefficient of thevolume of the object, wherein the absorption coefficient is a functionof the first acoustic spectrum and the second acoustic spectrum.
 2. Themethod of claim 1 wherein quantifying the absorption coefficient of thevolume of the object comprises dividing the first acoustic spectrum bythe second acoustic spectrum frequency by frequency to generate anormalized acoustic spectrum, and fitting the normalized acousticspectrum to an absorption coefficient curve.
 3. The method of claim 1wherein the light source comprises a first light source for emitting thefirst light pulse and a second light source for emitting the secondlight pulse.
 4. The method of claim 1 wherein the first wavelength isdifferent from the second wavelength.
 5. The method of claim 1 furthercomprising disregarding a portion of the first acoustic spectrum and thesecond acoustic spectrum outside of a system bandwidth, wherein thesystem bandwidth is at least partially determined as a function of abandwidth of the acoustic transducer.
 6. The method of claim 1 furthercomprising determining at least one of HbT, HbO₂, HbR, and sO₂ values asa function of the quantified absorption coefficient.
 7. A system forquantifying an optical absorption coefficient of a volume of an object,said system comprising: a light source for emitting a first light pulsehaving a first wavelength and a second light pulse having a secondwavelength; an optical element for directing the first light pulse andthe second light pulse into a volume of the object; an acoustictransducer for receiving a first acoustic wave emitted by the object inresponse to the volume of the object absorbing the first light pulse anda second acoustic wave emitted by the object in response to the volumeof the object absorbing the second light pulse; and a controller forperforming a Fourier transformation on the received first acoustic waveand second acoustic wave to generate a first acoustic spectrum and asecond acoustic spectrum, respectively, and quantifying the absorptioncoefficient of the volume of the object as a function of the firstacoustic spectrum and the second acoustic spectrum.
 8. The system ofclaim 7 wherein the controller is further configured to quantify theabsorption coefficient of the volume of the object by dividing the firstacoustic spectrum by the second acoustic spectrum, frequency byfrequency, to generate a normalized acoustic spectrum and to fit thenormalized acoustic spectrum to an absorption coefficient curve.
 9. Thesystem of claim 7 wherein the light source comprises a first lightsource for emitting the first light pulse and a second light source foremitting the second light pulse, and wherein the first wavelength isdifferent from the second wavelength.
 10. The system of claim 7 whereinthe controller is further configured to disregard a portion of the firstacoustic spectrum and the second acoustic spectrum outside of a systembandwidth, wherein said system bandwidth is at least partiallydetermined as a function of a bandwidth of the acoustic transducer. 11.The system of claim 7 wherein the controller is configured to quantifyat least one of HbT, HbO₂, HbR, and sO₂ values as a function of thequantified absorption coefficient.
 12. A system for generating an imageof an object and quantifying an optical absorption coefficient of avolume of an object, said system comprising: a light source for emittinga first light pulse having a first wavelength and a second light pulsehaving a second wavelength; an optical element for directing the firstlight pulse and the second light pulse into a volume of a plurality ofvolumes of the object; an acoustic transducer for receiving a firstacoustic wave produced by the object in response to the volume of theobject absorbing the first light pulse and a second acoustic waveproduced by the object in response to the volume of the object absorbingthe second light pulse; a controller for performing a Fouriertransformation on the received first acoustic wave and second acousticwave to generate a first acoustic spectrum and a second acousticspectrum, respectively, quantifying the absorption coefficient of thevolume of the object as a function of the first acoustic spectrum andthe second acoustic spectrum and for reconstructing an image fromacoustic spectra detected from the plurality of volumes of the object;and a scanner for altering a spatial relationship between the object andthe optical element after the first acoustic wave and the secondacoustic wave are detected by the acoustic transducer for the volumesuch that the first light pulse and the second light pulse are directedinto another volume of the plurality of volumes of the object.
 13. Thesystem of claim 12 wherein the controller is configured to quantify theabsorption coefficient of the volume of the object by dividing the firstacoustic spectrum by the second acoustic spectrum, frequency byfrequency, to generate a normalized acoustic spectrum and fits thenormalized acoustic spectrum to an absorption coefficient curve.
 14. Thesystem of claim 12 wherein the light source comprises a first lightsource for emitting the first light pulse and a second light source foremitting the second light pulse, and wherein the first wavelength isdifferent from the second wavelength.
 15. The system of claim 12 whereinthe controller is configured to disregard a portion of the firstacoustic spectrum and the second acoustic spectrum outside of a systembandwidth, wherein said system bandwidth is at least partiallydetermined as a function of a bandwidth of the transducer.
 16. Thesystem of claim 12 wherein the controller is configured to quantify atleast one of HbT, HbO₂, HbR, and sO₂ values as a function of thequantified absorption coefficient.